Scintillator materials are used to detect γ-rays, x-rays, neutrons and electrons in research and medical imaging devices. Furthermore, high-energy and nuclear physics relies on scintillation to detect weakly interacting particles and energies such as dark matter and dark energy. Some of the basic requirements for scintillator materials are: (1) a fast response time in the range of 10-100 ns for time-resolution, (2) a high light yield in excess of tens of thousands of photons per absorbed radiation particle, (3) a high density p and atomic number for efficient y-ray detection, since the latter is proportional to ρZ3-4, (4) good match of the scintillation output with the sensitivity of light sensor (i.e. photomultiplier tube), and (5) availability of large single crystals. Early scintillation materials to detect radiation were K2Pt(CN)4 and CaWO4 introduced in 1895 by Roentgen and in 1896 by Becquerel. CaWO4 converts x-rays into blue light and was used early on for x-ray photography and medical imaging. The light yield (Yhy) of scintillators given in photons/MeV can be approximated by a simple formula: Yhv=[106/βEg] S QE, with β being a materials dependent constant (2.5 for halides), Eg the band gap of the host lattice, S the energy transfer efficiency from the host lattice to the activator and QE the quantum efficiency of the activator. The product of βEg is the energy required to produce one exciton. To maximize light yield S and QE must be close to 100% and Eg small.
The two main techniques used in medical imaging are X-ray tomography and positron emission tomography (PET). In X-ray tomography, radioisotopes are injected into the body by administering compounds containing radioactive isotope such as 99Tc. The emitted radiation in the 120-150 keV energy range is then detected using a scintillator material. By appropriate camera rotations around the patient and the use of multiple detectors mathematical algorithms are used to reconstruct a three-dimensional image of the biological entity investigated. Modern computer tomography (CT) scanners use thousands of individual X-ray detectors spaced about 1 mm apart. Scintillators used for CT are (Y,Gd)2O3:Eu3+, Gd2O2S:Pr3+and CdWO4. The emissions are at 612 nm due to a 5D0→7F2 transition in (Y,Gd)2O3:Eu3+and 510 nm due to a 3P0→3H2, 3F3 transition in Gd2O2S:Pr3+. The 480 nm emission in CdWO4 is a charge transfer transition. The various tungstates AWO4 (A=Ca, Ba, Cd, Zn) adopting the scheelite structure are also widely used scintillator materials. Again a strongly distorted excited state that differs significantly from the ground state leads to the emission of a broad-band with a large Stokes shift. The Cd2+4d electronic states are located near the bottom of the valence band which is formed by the oxygen 2p orbitals. The conduction band has mainly W 5d character. The scintillator properties of this material are based on the WO42− entities and can be rationalized as resulting from the charge transfer out of the O2−2p states into W6+5d0 states with contributions from Cd2+4d states. In these tungstates thermal quenching at room temperature is usually very small and quantum efficiencies of up to 70% can be achieved. Using a simple approximation for the energy efficiency of scintillator materials developed above a theoretical conversion efficiency of 6% can be calculated. Experimental values of 3.5% have been found. The light yields of all these scintillators are moderate and produce about 10,000 photons/MeV.
In PET, the annihilation of positrons is exploited for imaging purposes. The predominantly used positron emitters are the isotopes of 11C (t1/2˜20 min), 13N (t1/2˜10 min, 15O (t1/2˜2 min) and 18F (t1/2˜110 min). Due to their positive charge and strong interaction with matter, the emitted positrons are stopped in biological tissue after traveling just a few millimeters. When slowing down, positrons will annihilate with electrons in condensed matter and emit in most cases two γ-rays in opposite directions which both have energies of 511 keV. PET makes use of this collinear emission of two γ-rays by measuring the temporal coincidence data along straight lines. This also permits the reconstruction of 3-dimensional biological objects using appropriate algorithms. Bi4Ge3O12(BGO) crystals are used in PET scanners as scintillator materials. The structure of Bi4Ge3O12 consists of isolated GeO4 tetrahedrons and Bi3+ ions which have an asymmetric coordination with three short (2.16 Å) and three long (2.60 Å) Bi-oxygen distances as a consequence of the 6s2 lone pair electronic configuration. In the excited state, this coordination is more symmetric. However, this material has high thermal quenching and at room temperature about ⅔ of the light efficiency of BGO is quenched. The experimental value of the energy efficiency of BGO is about 2%. The density of BGO is 7.1 g/cm3 and its effective Z with 75 is very high. The Stokes shift is quite large with 14,000 cm−1 which minimizes self-absorption and allows thin slabs to be used as detectors since the crystal is transparent to its emission light at 480 nm.
Nal:Tl+ is another commonly scintillator material used in PET and as an x-ray phosphor. Its density is about half of that of Bi4Ge3O12(3.86 g/cm3) and its Zeff is 51. With a light yield of about 40,000 photons/MeV , a decay time of 230 ns, non-proportionality of the light yield in the 60-1275 keV range and hygroscopic behavior requiring the crystals to be sealed one would not give this material a big chance for market penetration. However, easy and low cost manufacturing have provided economic opportunities despite rather mediocre technical specifications.
As such, a need exists for improved scintillator materials.